Two general imaging problems in radiology involve the determination of a radiation source distribution and/or the effect of a filter, in this case a patient, on the radiation source distribution. Consider the field of nuclear medicine where the radiation source or other radionuclide distribution emits photons or positrons, Image data acquisition in nuclear medicine presents several challenges in addition to constraints imposed by finite acquisition times and patient exposure restrictions. Most photon energies that are of interest in nuclear medicine are higher than the typical photon energies employed in diagnostic x-ray radiography. In particular, Positron Emission Tomography (PET) involves the detection of pairs of very high energy photons due to annihilation events Unfortunately, the photon radiation source, such as a radionuclide, used in nuclear medicine is not directional and the source distribution within the body is not precisely known.
Photons that escape the body may be scattered, altering their energies and/or direction vectors. It is desirable for many applications to discriminate against scatter radiation reaching the detector based on energy and/or direction. It may also be desirable to only detect radiation with a specific direction vector, since many detection systems possess poor directional discrimination capability and have finite response times within which to detect events, thereby limiting detection rates. Thus detection systems used in nuclear medicine such as Gamma cameras or PET scanners often employ conventional, such as attenuating or rigid geometry, focused or unfocused collimators, often referred to as grids or grid collimators, to help define the direction vectors of a detected photons. The direction vectors and energies of non-scattered photons are well-defined. Unfortunately, the emission of photons from the source distribution is non-directional and the radiation source distribution itself is typically not well-defined. A Compton-scattered photon suffers an energy loss and change in direction vector whereas a coherent or Rayleigh scattered photon only has its direction vector altered. In general x-ray radiography the source is a x-ray tube, although a radionuclide maybe substituted, used in a point, slit, slot, or area imaging configuration. The energy distribution and direction vector of the radiation from a x-ray tube are approximately known. These parameters are typically well-defined for a collimated radionuclide source used in an application such as point-scan Compton scatter imaging and material analysis. A number of detection formats are in use depending on the application. A planar detector geometry is typically utilized for applications such as mammography, angiography, and chest radiography which typically employ detectors such as x-ray film-screen devices, or storage phosphor screens, or image intensifiers coupled to cameras. Slit- and slot-scan formats are also available, usually incorporating improvements to the detectors and, in some instances, the radiation source. Additional image acquisition formats include ring-shaped detectors or flat detectors for fan-beam or cone-beam tomography, respectively. Common detector geometries used in nuclear medicine typically include one or more planar detectors, which are basically standard Gamma cameras, with attached conventional collimators or ring detectors, used in Positron Emission Tomography. Imaging systems based on standard Gamma camera and related detector designs are frequently used for a number of nuclear medicine studies such as heart, brain, thyroid, gastro-intestinal, whole body, and breast imaging, including scintimammography. A basic Gamma camera design employs a large, planar array of scintillation crystals or a single, large, planar scintillation crystal optically coupled to an array of photomultiplier tubes (PMTs). A conventional focused or unfocused collimator is typically mounted to the face of the Gamma camera. This inflexible imaging system is then positioned such that the region of interest containing the source distribution is within the field of view. It provides a limited degree of spatial resolution and energy resolution while removing some fraction of unscattered and scattered radiation that would otherwise degrade image quality. Unfortunately a substantial fraction of useful unscattered radiation is also attenuated. Another infrequently used design replaces the conventional collimator with a coded aperture such as a uniformly redundant array aperture which is also based on photon attenuation and is typically rigid. Commercial systems may use one, two, or three Gamma camera detector units. One commercial system eliminates the use of scintillator crystals and PMTs with a rigid, planar, 2-D CdZnTe semiconductor detector manufactured by abutting four 2-D CdZnTe arrays of moderate size. Techniques for abutting 2-D silicon arrays are well-known in the art. Drawbacks to employing large- or medium-sized 2-D CdZeTe arrays capable of high detection efficiency include the difficulty of growing thick CdZnTe crystals with acceptable levels of defects and creating a low noise, 2-D array readout structure on top of a large- or medium-size CdZnTe crystal. Grid collimators are still desirable for many applications since the direction vectors of detected photons are otherwise poorly defined. A design which replaces a conventional collimator with a relatively thin, planar semiconductor, often Ge, array of moderate size, which serves as a Compton scatterer is referred to as a Compton electronic Gamma camera. This system is still being refined. The detector module array described below can be used in place of a standard Gamma camera in a Compton Gamma camera system.
Nuclear medicine imaging applications are complicated by the fact that the spatial distribution of the source within a region of the patient is poorly defined. One way to simplify this problem is to use emitted photons of known energies. For example, a source that has one or more emission energies of a narrow energy bandwidth may be utilized. The problem now is the reconstruction of the source distribution rather than the calibration of the source distribution. The measured source distribution, i.e., the apparent source distribution, represents the filtered true source distribution, assuming self-attenuation is small. In certain nuclear medicine applications estimates of the true source distribution are obtained by calibrating the contribution of the filter, which may be the patient, to the apparent source distribution. Photon transmission measurements are made in order to estimate the effect of tissue scattering and absorption or attenuation on radiation source measurements by using a reference source that is external to the patient. Unfortunately, measuring photon transmission through the body does not duplicate the actual imaging chain acquisition format used in nuclear medicine where photon are transmitted out of the body. Photons in the two instances do not traverse comparable paths.